High magnetic field electromagnets have become important elements in various types of equipment over recent years. One important type of such equipment is medical imaging equipment, such as the type commonly referred to as magnetic resonance imaging (MRI) equipment. This equipment utilizes the mechanism of nuclear magnetic resonance (NMR) to produce an image, and accordingly imaging systems operating according to this mechanism are also commonly referred to as NMR imaging systems.
As is well known in the field of MRI, a high DC magnetic field is generated to polarize the gyromagnetic atomic nuclei of interest (i.e., those atomic nuclei that have nonzero angular momentum, or nonzero magnetic moment) contained within the volume to be imaged in the subject. The magnitude of this DC magnetic field currently ranges from on the order of 0.15 Tesla to 2.0 Tesla; it is contemplated that larger fields, ranging as high as 4.0 to 6.0 Tesla, may be useful in the future, particularly to perform spectroscopy as well as tomography.
The volume of the subject to be imaged is defined by the application of a gradient magnetic field in combination with the DC field. As the gyromagnetic nuclei in the defined volume will have a common resonant frequency different from atoms outside of the volume, modulation of the gradient field allows sequential imaging of small volumes. The images from the small volumes are then used to form a composite image of the larger volume, such as the internal organ or region of interest.
Imaging is accomplished in MRI by using the nuclear magnetic resonance mechanism in gyromagnetic atomic nuclei in the subject. The MRI apparatus uses an oscillator coil to generate an oscillating magnetic field that is oriented at an angle relative to the DC field, and that has a frequency matching the resonant frequency of the atoms of interest in the selected volume. Frequencies of interest in modern MRI are in the radio frequency range. The MRI apparatus also includes a detecting coil in which a current can be induced by the nuclear magnetic dipoles in the volume being imaged.
In operation, as is well known, the magnetic dipole moments of those atoms in the volume which are both gyromagnetic and also resonant at the frequency of the oscillating field are rotated from their polarized orientation by the resonant RF oscillation by a known angle, for example 90.degree.. The RF excitation is then removed, and the induced current in the detecting coil is measured over time to determine a decay rate, which corresponds to the quantity of the atoms of interest in the volume being imaged. Incremental sequencing of the imaging process through the selected volume by modulations in the gradient field can provide a series of images of the subject that correspond to the composition of the subject. Conventional MRI has been successful in the imaging of soft tissues, such as internal organs and the like, which are transparent to X-rays.
The strength of the DC magnetic field determines many attributes of the MRI process and results therefrom. A higher DC magnetic field can be used to resonate the gyromagnetic nuclei of elements other than hydrogen, so that in-vivo spectroscopy may be performed (in addition to tomography, or imaging). Furthermore, it is also well known in the art that the spatial resolution of MRI tomography improves as the strength of the available magnetic field increases. It is therefore desirable to provide MRI equipment with extremely high DC magnetic fields, such as 1.5 Tesla or greater.
Superconducting coils have generally been required to produce such extremely large magnetic fields, due to the large ampere-turns necessary to generate such fields. The superconducting material and accompanying cryogenic systems in such magnets add significantly to the cost of the imaging equipment. In addition, the size and weight of the MRI apparatus generally increases with the DC field strength of the magnet, as the weight and size of the magnet will increase with its DC field strength. Some conventional MRI magnets are sufficiently heavy (e.g., on the order of twenty tons) as to limit the installation of the MRI apparatus to a basement or ground floor laboratory.
High field electromagnets also generate high fringe fields extending away from the bore of the magnet. These high fringe fields can cause upset or erroneous operation of electrical equipment located near the magnet. As a result, the electronic equipment must be located either in a separate room from the apparatus, or behind a ferromagnetic shield away from the apparatus (or both, in the case where the room walls include iron shielding).
In one type of conventional superconducting magnet, electrical shielding surrounds the magnet bore to confine the field into a closed magnetic flux loop, and thus reduce the fringe magnetic field. Such electrical shielding is implemented by way of outer coils surrounding the main field-generating coils that conduct current in the opposite direction from that of the main field generating coils. The reverse polarity in the shielding coils generates a magnetic field that opposes the main field generated by the inner set of coils at locations outside of the bore, thus reducing the fringe field. However, the field generated by the shielding coils also tends to weaken the main field within the bore, and reinforce the field in the space between the main coil and the shielding coils.
An example of a conventional superconducting magnet which relies substantially on active superconducting shielding loops is described in U.S. Pat. No. 4,595,899. The magnet disclosed in this reference has a set of three driving coils surrounded by three shielding coils, with the current through the shielding coils adjusted to exactly cancel the dipole outside of the magnet. This reference further discloses the use of external ferromagnetic shielding located around the shielding coils to assist in further shielding. In this magnet, however, the bulk of the shielding is disclosed as effected by the outer shielding coils, considering the precise calculation of the dipoles set up by the driving coils and by the shielding coils, by which the main and opposing fields were matched to effect good shielding.
As indicated in U.S. Pat. No. 4,595,899, and as is true for other conventional electrically shielded magnets, any ferromagnetic shielding used in the magnet is generally located some distance away from the magnet bore. Such placement is intended to limit the effect of iron on the shape and uniformity of the magnetic field in the bore, because, as is well known in the art, iron or other ferromagnetic material near the bore will non-linearly affect the field within the bore, especially at fields above the threshold of magnetic saturation for iron at about 1.0 to 1.3 Tesla. As a result, the sole effect of the iron in these conventional magnets is to provide fringe field shielding at some distance from the magnet, with minimal effect on the field within the bore intended. In some cases, the ferromagnetic shield is located as far away from the bore as to be within the walls of the room surrounding the magnet (or MRI apparatus containing the magnet). This distancing of the ferromagnetic material from the bore causes significant problems in use of the magnets and equipment, either requiring large "footprints" for the magnet and its shielding, or requiring the specially constructed rooms to house the magnet or NMR equipment, either approach resulting in high cost and poor space utilization.
Examples of other prior magnets used in MRI are described in U.S. Pat. No. 4,612,505, in which shielding is accomplished by way of magnetic soft iron rods, conducting coils, or both. In particular, FIG. 3 of U.S. Pat. No. 4,612,505 discloses the use of a pair of relatively large superconducting shielding coils disposed outside the magnet. In addition, FIG. 4 of this reference illustrates a magnet having a shielding sleeve of magnetic soft iron, and shielding coils disposed outside thereof. The magnets disclosed in this reference have relatively low field strengths, such as on the order of 0.25 to 0.3 Tesla, and somewhat high fringe fields, such as 10 gauss or greater at a distance of three meters from the magnet axis.
U.S. Pat. No. 5,012,217, issued Apr. 30, 1992, describes yet another prior superconducting magnet utilizing a combination of active and passive shielding. This reference discloses the placement of a passive ferromagnetic shield around the main driving solenoid, but within the shielding solenoid (which generates the opposing magnetic field). This construction apparently requires that the large mass of the ferromagnetic shield be placed within the cryostat, substantially increasing the cryogenic load and, accordingly, the cost of maintaining the superconducting coils at superconducting temperatures.
In conventional magnets utilizing electrical shielding, either alone or in combination with ferromagnetic shielding, the cost of superconducting material for the outer coils is on the same order as that for the inner driving coils. The cryogenic load is also quite large for superconducting actively shielded magnets, due to the additional superconductors. In addition, it is believed that it is difficult to achieve uniformity of the magnetic field within the bore of the magnet where shielding is accomplished by cancellation of opposing fields, particularly where the desired magnetic field is 1.5 Tesla or greater.
As such, other prior work has been done in the field of superferric shielded superconducting magnets, as described in U.S. Pat. No. 4,783,628 (issued Nov. 8, 1988) and in U.S. Pat. No. 4,822,772 (issued Apr. 18, 1989), both incorporated herein by this reference and commonly assigned with this application. The magnets described in these patents utilize passive shielding of ferromagnetic material, such as iron, configured in such a manner that the shielding is superferric. Superferric shielding refers to the use of ferromagnetic material which affects the field within the magnet bore, particularly to enhance the field. As described in these patents, these superferric shielding concepts are useful even for magnetic fields significantly larger than the saturation threshold for the ferromagnetic material at which non-linear effects begin (e.g., above 1.0 to 1.3 Tesla for iron), and up to larger field strengths on the order of 4 Tesla. The construction of the magnets described in these patents provide a highly efficient magnet, considering the magnetic field strength as a function of the current conducted in the superconducting loops. Furthermore, the shielding is accomplished in such a manner that the uniformity of the field in the magnet bore is very high, even at very strong magnetic fields such as on the order of 4 Tesla, and with minimal fringe field (5 gauss at 50 to 100 cm from the outer wall of the bore).
Additional discussion of the effect of iron on the field within the magnet bore is presented in Siebold, et al., "Performance and Results of a Computer Program for Optimizing Magnets with Iron", IEEE Trans, Magnetics, Vol. 24, No. 1 (IEEE, January 1988), pp. 419-422. As particularly noted in FIG. 3 of this article, the coil system must be designed and adapted relative to the iron yoke in order to provide a uniform field in the bore.
The weight and size of the superferric shielded magnets described in U.S. Patent Nos. 4,783,628 and 4,822,772 can be quite substantial, however, such as on the order of 35 to 130 tons (as compared with actively shielded magnets weighing on the order of 20 tons). As a result, when used in medical equipment such as NMR stations, the "footprint" required for installation of such a magnet, as well as the weight-bearing capability of the floor of the room, are both significant. It is, of course, desirable to reduce the physical size and weight of NMR equipment, thus reducing the cost of the NMR laboratory. Besides the large footprint of conventional NMR magnets, it has been observed that many patients are uncomfortable when placed in magnets of such length, as the patient's entire body is generally disposed within the magnet during much of the imaging procedure. Indeed, conventional cylindrical NMR magnets have been referred to as "tunnel" magnets, representing the sensation perceived by the human subject when placed inside for an imaging procedure. It is therefore also desirable to provide a high magnetic field magnet for purposes of NMR which has good field homogeneity, but where the axial length of the bore is as short as possible.
By way of further background, it should be noted that the driving coils for magnets such as described in the above-referenced U.S. Pat. Nos. 4,783,628 and 4,822,772 are cylindrical in shape, so that a uniform magnetic field is provided over a portion of the axial length of the bore. As described, for example, in U.S. Pat. Nos. 4,587,490 and 4,590,428, and in Everett, et al., "Spherical coils for uniform magnetic fields," J. Sci. Instrum., Vol. 43 (1966), pp. 470-74, it is also known to provide spherical or quasi-spherical coil arrangements to produce a homogeneous magnetic field within the bore.
In addition, it is also known to provide error, or trim, coils in conventional iron-shielded magnets to provide adjustment of the homogeneity of the magnetic field within the bore. One example of such a magnet is described in U.S. Pat. No. 4,490,675, in which the error coils are disclosed as being within the soft iron cylindrical shield. U.S. Pat. Nos. 4,590,428 and 4,587,490 also disclose NMR or MRI magnets including main and error coils within an iron cylinder.
By way of further background, U.S. Pat. No. 4,924,185 discloses another cylindrical superconducting magnet. As disclosed therein, the sense of oppression on the part of the patient is reduced as the ratio of bore length to bore diameter is below 1.90.
Copending application Ser. No. 715,552, filed Jun. 14, 1991, entitled "A Compact Shielded Superconducting Electromagnet", incorporated herein by this reference and commonly assigned herewith, describes another cylindrical superconducting magnet which advantageously uses a combination of superferric shielding outside of the shielding coils. The magnet disclosed therein thus can be shorter in length while still providing high DC field in the bore and low fringe field away from the magnet.
As noted above, an important use of MRI is in the medical context, where human patients are placed within the DC magnetic field for the imaging of internal organs. Particularly for patients who are seriously ill, it is essential that the magnet be able to receive the patient without requiring disconnecting life support or monitoring conduits from the patient. Medical personnel must also be able to access the patient during the procedure. Furthermore, many patients become anxious or otherwise uncomfortable when placed within conventional MRI equipment, particularly of the cylindrical type where the bore is on the order of 2 meters long, such as in the case of the conventional magnets described hereinabove.
Furthermore, as noted above, it is also highly desirable that the weight and "footprint" of the magnet be minimized so that the MRI apparatus may be located in conventional and reasonably sized laboratory space throughout the medical facility, and not limited to installation on a ground floor due to the weight of the apparatus. Furthermore, it is desirable that the fringe field be sufficiently low that electronic monitoring equipment and other instrumentation may be kept in the same room as the MRI apparatus, and without requiring special shielding in the walls or an excessively large room.
It is therefore an object of the present invention to provide an extremely compact superconducting magnet having a high degree of effective shielding.
It is a further object of the present invention to provide such a magnet having a large aperture into which a patient may be placed.
It is a further object of the present invention to provide such a magnet having relatively light weight and low cost.
It is a further object of the present invention to provide such a magnet which can be fabricated using a single cryostat.
Other objects and advantages of the present invention will be apparent to those of ordinary skill in the art having reference to the following specification together with the drawings.